Transaxial compression technique for sound velocity estimation

ABSTRACT

An improved ultrasonic pulse-echo method and apparatus that has particular application in estimating sound velocity in organic tissue is disclosed. The method employs a standard transducer or transducer containing device which is translated transaxially, thereby compressing or displacing a proximal region of a target body in small known increments. At each increment, a pulse is emitted and an echo sequence (A-line) is acquired from regions within the target along the sonic travel path or beam of the transducer. Segments of the echo sequence corresponding to a distal region within the target are selected as a reference to estimate the incremental change in echo arrival time. A plot of these arrival time estimates versus the target compression depth is then generated and a least squares linear fit is made. The slope of the linear fit is c -1 , where c is an estimate of the speed of sound in the target.

This invention was made with government support under GRANT CA 38515 andCA 44389, amended by the National Institutes of Health. The governmenthas certain rights in the invention.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates generally to methods and apparatus for performingultrasonic diagnosis of a target body. More particularly, the inventionpertains to methods and apparatus for the measurement of sound speed ina target body. The invention is especially concerned with techniques forenhancing the accuracy of sound velocity measurements in compressibletargets using one or more ultrasonic transducers in pulse-echo mode.

2. Description of Related Art

Traditional ultrasonic diagnosis is achieved by transmitting ultrasonicenergy into a target body and generating an image from the resultingecho signals to survey anatomical structures. A transducer is used toboth transmit the ultrasound energy and to receive the echo signals.During transmission, the transducer converts electrical energy intomechanical vibrations. Acquired echo signals produce mechanicaloscillations in the transducer which are reconverted to electricalsignals for amplification and recognition.

A plot or display (e.g., on an oscilloscope, etc.) of the electricalsignal amplitude vs. echo arrival time yields the amplitude line(A-line) or echo sequence corresponding to a particular ultrasonictransmission. When the A-line is displayed directly as a sinusoidalpattern modulating at radio frequency (RF) it is referred to as an RF or"undetected" signal. For imaging, the A-line is often demodulated to anon-RF or "detected" signal.

Ultrasound techniques have been extensively used in the field ofdiagnostic medicine as a non-invasive means of analyzing the propertiesof tissue in vivo (i.e., living). A human or animal body represents anon-homogenous medium for the propagation of ultrasound energy. Acousticimpedance changes at boundaries of varying density and/or sound speedwithin a target body. A portion of the incident ultrasonic beam isreflected at these boundaries. Inhomogeneities within the tissue formlower-level scatter sites that result in additional echo signals. Imagesmay be generated from this information by modulating the intensity ofpixels on a video display in proportion to the intensity of echosequence segments from corresponding points within the target body.

Conventional imaging techniques are widely used to evaluate variousdiseases within organic tissue. Imaging provides information concerningthe size, shape and position of soft tissue structures using theassumption that sound velocity within the target is constant.Qualitative tissue characterization is carried out by interpretation ofthe grey scale appearance of the echograms. Qualitative diagnosislargely depends on the skill and experience of the examiner as well assystem characteristics. However, images based only on relative tissuereflectivity cannot be used for a quantitative assessment of diseasestates.

Techniques for quantitative tissue characterization using ultrasound areneeded for more accurate diagnosis of disease. One of the most promisingparameters for quantitative measurement is sound speed. Speed of soundchanges within regions of varying density and/or molecularcompressibility within the tissue. Thus, it is expected that changes intissue density due to disease will result in changes in the speed ofsound. Indeed, it has been shown that changes in the speed of sound intissue often correlate with tissue pathology. For example, cirrhoticliver tissue has been observed to contain more fat than normal livertissue. The velocity of sound in cirrhotic tissue would therefore beexpected to be lower than in normal tissue. Similarly, changes in tissuedensity in the region of tumors may result in changes in sound velocityin the tumor region. Unfortunately, however, such changes are relativelysmall and account for up to only 10% of the speed of sound in normaltissue. Therefore, accuracy in sound velocity estimation is extremelyimportant in the analysis of tissue for pathological conditions.Usually, the accuracy of sound velocity estimations must be at least1.0% to have specific value for quantitative tissue characterization.Hence, a need exists for the accurate measurement of sound velocity inorganic tissue for clinical diagnosis.

Traditionally, measurement of sound speed has been conducted withtransmission techniques. A first method of sound velocity measurementinvolves the transmission of sound pulses through tissue regions ofknown dimension and recording the time required for the pulse totraverse the region. The quotient of travel distance and travel time iscomputed to yield the velocity. However, due to the softness of mosttissues, the dimensions of the tissue sample cannot be accuratelymeasured which results in an error-prone measurement of sound velocity.Moreover, a reference liquid with a known speed of sound may be requiredto calibrate the apparatus.

A second transmission technique that has been used in medicaldiagnostics involves a transmitting transducer and a separate receivingtransducer arranged so that they are aimed at one another with theirrespective axes of radiation coincident. The body of the subject isplaced between the transmitting and receiving transducers. However, invivo application of this technique has been limited to accessible organslike the breast or testes; other in vivo applications can be adverselyaffected by such factors as bowel gas, bone and inaccessibility.

A third transmission technique is disclosed by Ophir and Lin, "ACalibration-Free Method for Measurement of Sound Speed in BiologicalTissue Samples", IEEE Transactions on Ultrasonics, FerroeIectrics, andFrequency Control, Vol. 35, No. 5, (1988) 573-577. This method allowsaccurate measurement of the speed of sound in soft tissue samples, whileovercoming the limitations of initial techniques. The method employs areceiving hydrophone and a transmitting transducer that are coaxiallyaligned opposite each other. The transmitting transducer is in contactwith the tissue sample, while the hydrophone penetrates the tissuesample at well-controlled incremental depths. The transit times of thepulse are recorded for all penetration depths of the hydrophone. Thesetransit times are then plotted against the relative depths of thehydrophone, and a linear regression fit is made to the data. The slopeof the fitted line is c⁻¹, where c is the estimated speed of sound inthe tissue sample. The technique requires neither calibration involvinga reference medium, nor the knowledge of the thickness of the tissuesample. Yet, while this technique is capable of accurate measurements oftissue in vitro, it is clearly not suitable for speed of soundestimations in vivo.

Several techniques have been proposed for the measurement of soundvelocity in vivo using ultrasonic transducers in pulse-echo mode. In onemethod, sound speed is measured using misregistration between pulse-echoimages of the same structure obtained with two different sound beams.Sound velocity is determined from the difference in position of the samefeature in different images. This method works best when a well-definedfeature is available. In simulated tissue regions, known as "phantoms",thin wire added to the region will provide such a well-defined feature.However, well defined features are not easy to find in living tissue andthe resulting sound speed measurement is therefore not as accurate. SeeRobinson et al., "Measurement of Velocity of Propagation from UltrasonicPulse-Echo Data", Ultrasound in Med. & Biol., Vol. 8, No. 4, (1982)413-320.

In another pulse-echo technique called the "focus adjustment method",the mean sound speed between a reflector and linear array transducer ismeasured using the following three parameters: time of flight, time offlight difference, and distance between two receiver elements. To detecttime of flight, the system delay-line time compensator is adjusted toobtain the sharpest reflector image. Thus, the sharpness of the targetis maximized by interactive user control of signal delays at thetransducer aperture. However, irregular tissue structures cause randomrefractions of the ultrasonic beams and make sharp focusing difficult.Also, the method is highly dependent on qualitative judgment. SeeHayashi et al., "A New Method of Measuring In vivo Sound Speed in theReflection Mode", J. Clin. Ultrasound, Vol. 16, (1988) 87-93.

A third pulse-echo method described in U.S. Pat. No. 4,669,482, involvesin vivo sound velocity estimation by identifying segments of differentsound velocity along a tracked ultrasonic beam using at least twowidely-separated acoustic vantage points. The tracked beam ispartitioned into at least two contiguous segments, the boundary betweenthe two segments being the inner body of the body wall fat. A pluralityof ultrasound pulse travel time measurements are made, each with adifferent apparent angle of intersection between the tracked beam andthe tracking beam. For each measurement, techniques are employed forcorrecting refraction occurring in a transverse plane. Data pairscollected in the plurality of measurements are fitted to an appropriateequation using curve-fitting techniques well known in the art, by whichthe index of refraction at the body wall inner boundary, the inclinationof the inner boundary, and the speed of sound in the internal tissue arederived. This technique, however, is not desirable in clinical settingsbecause of the large "footprint" of the apparatus on the patient thatresults in a cumbersome examination procedure. Also, inaccuracies due tobone and/or bowel gases are common because of the wide spacing betweentransmitting and receiving transducers.

Hence, all the above pulse-echo techniques are clinically limited due tothe need to use two widely separated acoustic vantage points and/or bythe requirement that an identifiable, discrete target be available inthe tissue. The use of two widely separated vantage points makes theapparatus and the examination procedure cumbersome, while the existenceof a discrete target cannot always be guaranteed. Another potentialproblem is due to the effects of the overlying fat layer of the body onthe estimation.

SUMMARY OF THE INVENTION

The present invention provides an improved pulse-echo method andapparatus that has particular application in estimating sound velocityin organic tissue. The present invention addresses the problems of priorpulse-echo techniques by providing a relatively small footprint andobviating the need for a readily identifiable, discrete target withinthe tissue.

According to the present invention, a standard transducer or transducercontaining device is translated transaxially, thereby compressing ordisplacing a proximal region of a target body in small known increments.At each increment, a pulse is emitted and an echo sequence (A-line) isacquired from regions within the target along the sonic travel path orbeam of the transducer. Segments of the echo sequence corresponding to adistal region within the target are selected as a reference to estimatethe incremental change in echo arrival time. A plot of these arrivaltime estimates versus the target compression depth is then generated anda least squares linear fit is made. The slope of the linear fit is c⁻¹,where c is an estimate of the speed of sound in the tissue.

The present invention takes advantage of the acoustical properties ofphysically compressible or displaceable materials. These materials oftencontain a large number of acoustic "scatterers." These scatterers, beingsmall compared to the wavelength of the sound frequencies involved, tendto reflect incident sound energy in all directions. For example, inhomogeneous tissue regions, the scatterers may comprise a collection ofnearly identical reticulated cells. The combined reflections from eachscatterer create a background echo signal called speckle. The presentinvention employs standard pattern matching techniques to track areference echo sequence segment corresponding either to a reflector orother echo source, such as speckle, in a distal tissue region within thetarget body. See, e.g., J. S. Bandat and A. G. Piersol, "Random Data:Analysis and Measurement Procedures," Wiley Interscience, New York 1971,pp. 30-31. A discrete reflector, like a bone or blood vessel, may beused as a reference if desired, but is not necessary; any arbitrarysegment of the backscattered echo sequence may be used as a reference.

Bias occasioned by distal deformation of the reference echo source dueto the proximal compression or displacement of the target may becorrected by using a second stationary transducer. The second transduceris oriented such that its beam intersects the beam of the firsttransducer at a small angle within the region of the reference echosource. The echo time delay due to the distal deformation is detected bythe second transducer and is used to unbias the sound velocity estimate.While two acoustic vantage points are used, they are maintained at closeproximity to each other, so that the total transducer "footprint" on thetarget is no larger than that which is due to a standard transducerarray.

The present invention is of particular interest in interrogating organictissue, especially human and other animal tissue. A principal object ofsuch interrogation is to detect echo signals in the tissue that maysuggest the presence of abnormalities. More specifically, the effect ofcompression or displacement of the tissue on the characteristics of theecho signals becomes a possible key to such detection. It will be notedat this point that the invention is contemplated to have significantapplications other than in the study of tissue. One such application,for example, may be materials and products such as cheese or crude oilthat are physically compressible or displaceable by movement of atransducer. Thus, as a transducer is pressed against such a material,particles within the material are displaced from one position toanother. For elastic materials, release of the pressure enables theparticles to return to their original position.

It will be noted that the transducers employed in the present inventionneed not be in direct contact with the materials to which they areapplied. It is necessary, however, that transducers be sonically coupledto the materials. Sonic coupling methods and agents are well known inthe art.

It will also be noted that a material may be interrogated according tothe invention either (a) by advancing a transducer against a material toincrease compression, or (2) by retracting a transducer from acompressed position within the material.

As noted above, it is not necessary that an echo from a discrete featurein a tissue or other compressible material be employed. It is sufficientthat an identifiable echo segment be present in the echo signalresulting from a transmittal signal. Even though the physical featurewithin a material responsible for a selected echo sequence segment maynot be clearly known, the selected echo segment is an adequate referencefor the purposes of the invention. Thus, compression of the material andthe signal travel times determined before and after such compression maybe based on such echo segments.

As stated above, the invention may be practiced either by compressing atransducer against a compressible material from an initiallynon-compressed condition, or by retracting a transducer from aninitially compressed condition. In either case, however, it ispreferable that the distance traveled by the transducer be less than thewavelength of the ultrasonic signal produced or received by thetransducer.

The present invention may also be employed for localized estimation ofsound speed in targets having multiple layers. The speed of sound ineach of progressively deeper layers is sequentially estimated byemploying the same techniques discussed above. Distal regions at layerboundaries are used as the echo source for arrival time estimates.According to the present invention, the speed of sound can be estimatedin each layer from only two echo sequences along the axis of radiation.Thus, imaging of the speed of sound parameter in a plane or volume of atarget body can also be accomplished by appropriate lateral translationof the transducers.

Other objects and advantages of the invention will become readilyapparent from the ensuing description.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1a shows an embodiment of the invention where one transducer issonically coupled to a target body to interrogate a distal tissue regionwithin the target body;

FIG. 1b shows a plot of the RF echo signal originating from the distaltissue region interrogated in FIG. 1a;

FIG. 2a shows the transducer of FIG. 1a imparting a small compression toa proximal region of the target body;

FIG. 2b shows a plot of the time shifted RF echo signal originating fromthe distal tissue region interrogated in FIG. 2a;

FIG. 3a shows the transducer of FIG. 1a imparting a further compressionto a proximal region of the target body;

FIG. 3b shows a plot of the further time shifted RF signal from thetissue region interrogated in FIG. 3a;

FIG. 4a shows an embodiment of the invention where both compressing andnoncompressing transducers are acoustically coupled to a target body tointerrogate a distal tissue region within the target body;

FIG. 4b shows a plot of the RF signal originating from the distal tissueregion interrogated in FIG. 4a from the vantage point of thenoncompressing transducer;

FIG. 4c shows a plot of a time shifted RF signal originating from thedistal region interrogated in FIG. 4a; from the vantage point of thenoncompressing transducer;

FIG. 5 shows an embodiment of the invention where two transducers areused to interrogate multiple tissue layers;

FIG. 6 shows an embodiment in which a transducer is sonically coupled toa target via a stand-off device containing an acoustic coupling fluid;and

FIG. 7 is a plot comparing corrected and uncorrected speed of soundestimations in simulated tissue.

DETAILED DESCRIPTION

The basic method resembles the penetrating hydrophone transmissiontechnique discussed above. An adaptation of this technique to thepulse-echo mode is used. A transducer is positioned on or otherwisecoupled to a target body and advanced axially toward the target in smallknown increments. As noted earlier, the invention may also be practicedby incrementally retracting a transducer from a previously compressedposition. Since the relatively large aperture size precludes penetrationof the tissue, small tissue compressions occur instead. At eachincrement, a pulse is emitted and echo sequence (A-line) segments fromone or more selected distal tissue regions are used as a reference. Anyarbitrary segment of the backscattered RF echo signal from within thetissue may be identified and used as a reference. The selected segment--wavelet--of the RF signal corresponds to a particular echo sourcewithin the tissue along the beam axis of the transducer. As thetransducer compresses the tissue, it moves closer to the echo source,thereby shortening the travel path of the pulse and corresponding echo.The change in arrival times for echoes originating from the echo sourceas the transducer is incrementally advanced (or retracted) is related tothe speed of sound in the tissue. Thus, the speed of sound may bedetermined even though the distance between the transducer aperture andthe selected echo source are unknown.

The present invention contemplates transducers that may bepiezoelectric, ferroelectric or magnetostrictive in nature. The presentinvention is not limited by the size, focusing properties or bandwidthof the transducer to be employed.

FIG. 1a shows the transducer 10 sonically coupled to a target body 15.An ultrasonic pulse 18 is shown propagating within beam 20 toward a echosource 25 on beam axis 12. As the pulse 18 propagates through the target15, corresponding echoes are generated and arrival times noted at thetransducer aperture 11. The combination of all echoes generated fromreflections within the beam 20 is the echo sequence or A-linecorresponding to pulse 18. A radio frequency ("RF") signal plot of theA-line acquired from pulse 18 is shown in FIG. 1b. The amplitude of thesignal in millivolts is plotted against echo arrival times inmicroseconds (μs). Latter arrival times correspond to progressivelydeeper regions within the target body 15. An echo wavelet 30, within achosen arrival time window, is selected as a reference. The time windowmay be selected based on anatomical data from ultrasound imaging, or maybe arbitrary, e.g., every x micro seconds. The wavelet 30 originatesfrom the echo source 25 that is at an unknown distance from thetransducer aperture 11.

FIG. 2a shows the transducer 10 being translated along axis 12 to imparta small compression (y₁) to the tissue. Alternatively, as shown in FIG.6, a transducer 80 may be associated with a stand-off device 85 whichallows the transducer 80 to be acoustically or sonically coupled to thetarget body 90 without being in direct contact with the target body. Inthis case the stand-off 85, and not the transducer, compresses thetarget. In either case, however, the incremental compressions of thetransducer or transducer containing device are dependent on thefrequency of the transducer employed. More specifically, the magnitudeof the incremental compressions are based on the wavelength which is afunction of transducer frequency. In general, incremental shifts of lessthan about one wavelength are employed unless a discrete target is usedas a reference. Otherwise, tracking the reference signal segment will becomplicated by phase wrap. For example, in ophthalmic diagnosis atransducer of about 20 mHz may be employed, whereas a transducer of 3-5mHz would be suitable for interrogating abdominal tissue. When atransducer of 3-5 mHz is used, the compressions are generally on theorder of several mm, preferably between 0.1-2 mm.

After the transducer 10 compresses the target, a second pulse 22 isemitted and the corresponding A-line segment is acquired from a desireddepth within the tissue. FIG. 2b shows the RF plot of a time shiftedA-line corresponding to pulse 22. The wavelet segment or 32 associatedwith echo source 25 is also time shifted. The time shifted wavelet 32 istracked within the selected time window using standard pattern matchingtechniques. The arrival time of wavelet 32 is prior to that of wavelet30 above, since the distance between aperture 11 and feature 25 wasshortened by the compression Y₂.

FIG. 3a shows further tissue compression (y₁) and a third pulse 24emitted after the compression. The RF plot of the A-line in FIG. 3bshows an additional time shift in the signal. The wavelet 35 is trackedwithin the selected time window and is used to note the signal timeshift. Assuming uniform sound speed and no displacement of the echosource involved in producing the RF signal wavelet of interest, thesound speed estimate in the tissue contained between the transducer andthe location of these scatterers is: ##EQU1## where n is the number ofuniform transducer compressional displacements, y_(i) is the ithcompression, and t_(i) is the ith measured temporal shift in thereference echo signal wavelet. The factor of 2 in the numerator accountsfor the pulse-echo nature of the technique in which ultrasound (pulses)travels to and returns (echoes) from the echo source in the selecteddistal region. However, the method of the present invention is notlimited to a particular algorithm for calculating the sound speedcharacteristics of a target body.

According to the present invention, the one transducer embodimentdiscussed above may be conveniently employed in instances in which thetarget body being interrogated contains very compressible materials.Also, the method may be adapted to compress the tissue and acquire anA-line segment prior to the arrival of an elastic wave associated withthe proximal compression. This is possible because, although the elasticwave travels at about 20 meters per second (m/s), the ultrasonic pulsetravels at about 1540 m/s. Thus, the A-line is obtained from theselected time window prior to the arrival of the elastic wave. However,this is not feasible in some instances. In these cases, the assumptionof no distal feature displacement is inadequate. Although thedisplacements of echo sources within the target will generally fall offasymptotically with range, minute displacements may occasionally bedetected even far from the transducer. When this occurs, it is necessaryto make a correction for the distal displacements.

To correct the estimate, the expression of eq. (1) is modified toreflect the presence of additional, unknown time delays t_(d),i due tosuch displacements indicated by the subscript d. Therefore, theresulting modified estimate of the speed of sound is: ##EQU2## Since thequantities (t_(i) -t_(d),i)≦ t_(i) are the actual time delays that aremeasurable, the estimate is always positively biased unless the t_(d),i=0.

Fortunately, the quantities t_(d),i can be independently estimated usinga second transducer. This is shown in FIG. 4a. In addition to thecompressing transducer 38, a stationary noncompressing transducer 40 isused, whose beam axis 42 is directed such that it intersects the beamaxis 52 of the compressing transducer 38 at the range that correspondsto the echo source 50. The noncompressing transducer 40 operates in thepulse-echo mode and detects minute displacements of the echo source 50in the region of beam intersection that appear at time shifts δt_(d),i.In a preferred embodiment, a pair of matched ultrasonic transducers isused for the compressing and noncompressing functions, respectively.Still, the invention also contemplates unmatched transducers, or thecombination of a compressing transducer and a noncompressing steerabletransducer array.

The beam of the noncompressing transducer may be oriented to intersectthe beam of the compressing transducer in the region of the echo sourceof interest by using known beam tracking techniques. For example, thecompressing transducer 38 may operate in pulse-echo mode and acquireecho sequence segments having a unique arrival time from echo source 50.The noncompressing transducer 40 can be spaced laterally from the firsttransducer to operate as a echo receiver. The noncompressing transducer40 is angled until an echo burst coincident in arrival time with thedesired echo source is received. The angle of orientation 44 is notedwhen the beams of both transducers intersect at the desired echo source50.

Continuing in FIG. 4a, the noncompressing transducer 40 emits a pulse 4sfrom the surface of the target 45. As pulse 48 travels through tissueregions at the intersection of beams S4 and 47, an RF echo wavelet 60(FIG. 4b) corresponding to an echo source at axis position 50 isreceived by the noncompressing transducer 40. Meanwhile, transducer 38is energized to emit a pulse and receive a corresponding echo fromsource 50. When transducer 38 is compressed a distance y_(i), an elasticwave 55 travels through the tissue and diminishes asymptotically withrange. The reference echo source 50 is moved slightly to position 51along the beam axis 52. Second pulses are emitted from both transducers38 and 40. A time shifted RF echo wavelet 62 (FIG. 4c) is received atnoncompressing transducer 40, since the echo source 51 is now furtherfrom aperture 49.

If the angle between the beams is θ₁, then

    Δt.sub.d,i =δt.sub.d,i cosθ.sub.1.       (3)

The values of δt_(d),i so obtained are added to the denominator of Eq.(2) to result in an unbiased corrected estimator: ##EQU3##

The method described so far estimates the speed of sound in a targetbody having a monolayer, which extends from the transducer aperture tothe depth of interest. The method of the present invention may beextended to allow local estimation of sound speed in layered media,where each layer may, in general, have a different sound speed. Theprocedure involves sequential estimation of the speed of sound inprogressively deeper layers.

Turning now to FIG. 5, the estimate of the speed of sound in the firstlayer 60, c_(ul), is determined by applying the basic procedure to thislayer whereby the boundary between the layers 67 is used as the echofeature. The speed of sound c_(u2) in the second layer 62 is determinedby aiming transducers 65 and 70 at an angle 72 of θ₂ at an echo source75 in the second layer 62. Selection of the echo source is based uponthe same criteria employed in the basic procedure outlined above. Theundesired temporal shifts at 67 of the distal margin of the first layer60 now become the forcing functions for the second layer 62, i.e., thespatial compressions imparted on the second layer are

    Δy.sub.2,i =c.sub.u1 δt.sub.d1,i cosθ.sub.1 (5a)

Advantageously, the quantities of y₂,i may be estimated as

    Δy.sub.2,i =Δy.sub.2,i -c.sub.u,1 (Δt.sub.1,i -Δt.sub.1d,i)                                       (5b)

Again, the minute displacements of the echo source 75 due to elasticwave 66 in the region of beam intersection appear as time shiftsδ_(d2),i cosθ₂ in one acquired RF signal plot. Applying equations (4)and (5a) to second layer 62 and temporarily assuming no displacement ofthe boundary 73 interrogated by the beam of the noncompressingtransducer as shown in FIG. 5, we get: ##EQU4##

The quantities (t₂,1 - t_(d2),i) in eq. (6) are now measured as the timeshift of the echo feature in the second layer before and after eachcompression, with respect to the boundary echo.

A slight complication arises in the approach of eqs. (6) and (7) if theregion of the boundary 73 between first and second layers 60 and 62 thatintersects the noncompressing beam is compressed as well. The assumptionbeing eq. (6) has been that such compression does not occur. If suchcompression occurs and if it is of the same magnitude as the compressionat 67 seen by the compressing transducer, then no additional usefulinformation is gained by the noncompressing transducer over that whichis available from the compressing transducer. In general, however, thecompression of region 73 of the boundary will be non-zero, but less thanthe compression of the boundary region 67 that is under the compressingtransducer. We observe that additional time delays (or advances) will bedetected by the noncompressing transducer due to the motions of theboundary, where in a region of small boundary displacements δδy_(d1),ithe speed of sound changes from c_(u2) to c_(u1). These measurabledifferences in the arrival time of the echoes from within the secondlayer due to this boundary displacement are given as ##EQU5## and thequantity δδy_(d1),i can be estimated in a similar fashion to eq. (5a) as

    δδy.sub.d1,i =c.sub.u1 δδt.sub.d1,i (9)

where δδt_(d1),i is the measurable additional delay in the arrival timeof the boundary echo at the noncompressing transducer. Combining eq. (8)and (9) yields ##EQU6## Adding this term to eq. (6) we get ##EQU7##where t_(d2),i is given by eq. (10).

Since t_(d2),i =δt_(d1),i cosθ₂, the ability to solve eq. (11) relies onthe inequality

     t.sub.2,i -δδt.sub.d2,i ≠0.             (12)

This condition will hold true only if the boundary displacements seen byboth transducers are unequal. This can usually be accomplished byproperly separating the transducers. Given that the inequality of eq.(12) holds, equations (10) and (11) constitute a system of two equationswith two unknowns, C_(u1) and C_(u2) ; the speed of sound in the deeperlayer is estimated from knowledge of the speed of sound in the precedinglayer and from some of the measured echo time delays. The layeredregions are selected based upon the differential compressibility fromfront to back. Thus, in highly compressible tissue, the selected layerscan be relatively thin, e.g., about 1 cm. For tissue regions that arenot very compressible, the selected layers are thicker to help insurethat sufficient differential compressibility is obtained for meaningfulchanges in signal path.

Although the apparatus and method of this invention have been describedin relation to clinical diagnosis, this should be understood not to be alimiting factor on the utility of the invention. To the contrary, thepresent invention has utility in any area in which the speed of sound oforganic tissue may be desired. For example, the present invention may beused in forensics, tissue characterization studies, veterinary medicine,laboratory experiments, and industrial applications. Also, the presenttechniques may be employed to any materials that are capable of beingphysically compressed or displaced. That is, a material which isinternally displaceable in response to pressure applied to the material.

The various aspects of the invention will appear more specifically inthe following examples that are purely illustrative and should not beconstrued to limit the scope of the invention.

EXAMPLE 1

A water tank experiment was conducted to test the method in a singlelayer using a simulated tissue phantom. A 150 mm×150 mm×50 mm block offine reticulated polyester sponge was placed in a beaker and distilledwater was added to completely immerse it. The beaker was placed in adesiccator and laboratory vacuum (≈0.5 bar) was applied forapproximately 15 minutes. Thereafter, the beaker was submerged in a 60gal. distilled water tank and the sponge removed and placed on a 1/4 in.polished stainless steel reflector. The sponge was allowed to reachtemperature equilibrium of 37°±0.5° C. A reference value for the speedof sound in the sponge phantom was obtained using the method shown inFIGS. 1-3, with the difference that echoes from a steel plate at thebottom of the phantom were used as the reference wavelets.

To determine the speed of sound using the method of the presentinvention, the method shown in FIG. 4a was used. Matched 13 mm, 3.5 mHztransducers were used. The compressing transducer was moved into thesponge at 0.4 mm increments, 10 increments total, for a totalcompression of 4 mm. The noncompressing transducer was aimed at targetregions at several depths in the sponge, and echoes were recorded fromboth transducers at each incremental motion. Both the biased andunbiased speeds of sound estimates were calculated.

For the single-layer experiment, the reference value for the sponge andwater phantom was found to be 1256±1 ms⁻¹. FIG. 7 summarizes the resultsfor the single-layer experiment. Curve A shows the biased estimatedspeed of sound c_(b) in the region of interest extending to differentdepths. Clearly, estimates that are made in regions close to thetransducer are positively biased by up to 25 percent. This bias isexpected from eq. (2). The bias in the estimate diminishesasymptotically toward the correct value, such that at a delay of 240microseconds it amounts to about +2 percent. This is because theundesired distal compression of the region usually diminishes at greaterdepths in the sponge. Curve B shows the same data after corrections forthe undesirable distal compressions, i.e., the quantity c_(u) is within1 percent or less of the value determined by direct measurement at alldepths.

EXAMPLE 2

To test the ability of the present method to measure the speed of soundin an underlying tissue layer, a second foam phantom was constructedusing the same material as before, but saturated in a water andPolyethylene Glycol 600 solution to elevate the speed of sound. Forreference, the speeds of sound in the two phantoms were determined usingthe measurement technique described in Example 1. Next, the two phantomswere placed on a steel plate in water. To measure the sound speed in thesecond layer using the method of the present invention, the method shownin FIG. 5 was used. The noncompressing transducer was aimed at a targetregion in the second phantom, and the compressing transducer was moved 5mm toward the plate, thereby compressing both phantoms. The change inarrival times of the echoes from the target region were recorded fromboth transducers. Multiple repetitions of the experiment were done andthe average of the observations was used in the calculations.

In the two-layer experiment, the reference value for the speed of soundin the sponge in water and PEG-600 solution was 1589 ms⁻¹. A 5 mm motionof the compressing transducer resulted in a 1.08 mm translation of theboundary between the two phantoms, as observed with this transducer. Thetranslation of the same boundary, but along the beam of thenoncompressing transducer was 0.41 mm. The sound speed in the underlyingphantom estimated using equation (6) was 1601 ms⁻¹, or +0.75 percenthigher than the reference value.

Although the invention has been described with a certain degree ofparticularity, it is to be understood that the above description hasbeen only by way of example. Numerous other changes will be apparent tothose reading the specification without departing from the spirit andthe scope of the invention as claimed.

What is claimed is:
 1. A method for estimating sound velocity in atarget medium, comprising the steps of:(A) acoustically coupling a firstultrasonic transducer to the surface of the medium; (B) energizing thefirst transducer to transmit a first ultrasonic signal from the surfaceof the medium along a path into the medium; (C) detecting the arrivaltime of a first echo signal of the first transmitted signal at the firsttransducer from a distal position along the path within the mediumcorresponding to a reference echo segment of the first echo signal; (D)displacing a portion of the medium proximal the first transducersufficiency to change said arrival time while maintaining acousticcoupling between the first transducer and the medium; (E) energizing thefirst transducer; (F) detecting the changes arrival time; and (G)repeating steps (D) through (F), and determining the velocity of soundin the medium from the following equation: ##EQU8## wherein c is thevelocity of sound, n is the number of transducer displacement, ΔYi isthe transducer displacement during the ith compression, and Δt_(i) isthe ith measured change in arrival time of the echo signal.
 2. A methodof estimating the velocity of sound in a target body with two ultrasonictransducers, each having a separate radiation axis, comprising the stepsof:a) sonically coupling a first transducer to the surface of the body;b) selecting an echo source within said body on the radiation axis ofsaid first transducer; c) sonically coupling a second transducer on thesurface of the body and orienting it such that the radiation axis of thesecond transducer intersects the radiation axis of said first transducerat said echo source; d) detecting an arrival time at said firsttransducer for an echo originating from said echo source in response toa first ultrasonic pulse from said first transducer; e) detecting anarrival time at said second transducer for an echo originating from saidecho source in response to a first ultrasonic pulse from said secondtransducer; f) translating said first transducer a known distance alongits radiation axis toward said echo source by applying sufficient forceto the body to compress the body between the first transducer and theecho source; g) detecting an arrival time at said first transducer foran echo originating from said echo source in response to a secondultrasonic pulse from said first transducer; h) detecting the arrivaltime at said second transducer for an echo originating from said echosource in response to a second ultrasonic pulse from said secondtransducer; and i) calculating the estimated velocity of sound in saidtissue using the arrival times detected in steps d), e), g), and h),together with the translation distance of step f).
 3. The method ofclaim 2 wherein said target body includes multiple layers, whichcomprises the further steps of:j) selecting a first echo source at aboundary between a first body layer and a second body layer; k)sonically coupling said first and second transducers on the surface ofthe body such that their respective radiation axes intersect at saidfirst echo source; l) repeating steps d)-i) to determine the speed ofsound in said first layer; m) selecting a second echo source within saidsecond layer; n) reorienting said first and second transducers on thesurface of the tissue such that their respective radiation axesintersect at said second echo source; o) detecting an arrival time atsaid first transducer for an echo originating from said second echosource in response to a third ultrasonic pulse from said firsttransducer; p) detecting an arrival time at said second transducer foran echo originating from said second echo source in response to a thirdultrasonic pulse from said second transducer; q) translating said firsttransducer a known distance along its radiation axis toward said secondecho source by applying sufficient force to the body to compress thebody between the first transducer and the second echo source; r)detecting an arrival time at said first transducer for an echooriginating from said second echo source in response to a fourthultrasonic pulse from said first transducer; s) detecting an arrivaltime at said second transducer for an echo originating from said secondecho source in response to a fourth ultrasonic pulse from said secondtransducer; t) calculating the estimated velocity of sound of saidsecond layer using: (1) the arrival times measured in steps e) and h) asrepeated in step l) together with the speed of sound of the first layerdetermined in step l), and (2) the arrival times detected in steps o),p), r), and s), together with the translation distance of step q).
 4. Amethod of estimating the velocity of sound in a target body with onetransducer, said transducer having a rotation axis, comprising the stepsof:a) sonically coupling the transducer to the surface of the body; b)selecting an echo source within said body on the radiation axis of saidtransducer; c) detecting an arrival time at said transducer for an echooriginating from said echo source in response to a first ultrasonicpulse from said transducer; d) translating said transducer a knowndistance along its radiation axis towards said echo source by applyingsufficient force to the body to compress the body between the transducerand the echo source; e) detecting an arrival time at said transducer foran echo originating from said echo source in response to a secondultrasonic pulse from said transducer such that said echo originatesfrom said echo source prior to displacement of the echo source by acompression wave associated with the tissue compression in step d); andf) calculating the estimated velocity of sound in said body using thearrival times measured in steps c) and e), together with the translationdistance of step d).
 5. The method of claim 4 wherein said target bodyincludes multiple layers, which comprises the further steps of:g)selecting a first echo source at a boundary between a first body layerand a second body layer on the radiation axis of said transducer; h)repeating steps c)-f) to determine the speed of sound in said firstlayer; i) selecting a second echo source within said second body layeron the radiation axis of said transducer; j) detecting an arrival timeat said transducer for an echo originating from said second echo sourcein response to a third ultrasonic pulse from said transducer; k)translating said transducer a known distance along its radiation axistowards said second echo source by applying sufficient force to the bodyto compress the body between the transducer and the second echo source;l) detecting an arrival time at said transducer for an echo originatingfrom said first echo source in response to a fourth ultrasonic pulsefrom said transducer such that said echo originates from said first echosource after the arrival of a compression wave associated with thetissue compression of step k); m) detecting an arrival time at saidtransducer for an echo originating from said second echo source inresponse to said fourth ultrasonic pulse such that said echo originatesfrom said second echo source before the arrival of said compression waveassociated with the tissue compression of step k); n) calculating theestimated velocity of sound in said second layer using: (1) the arrivaltimes measured in step c) and e) as repeated in step h), together withthe speed of sound of the first layer determined in step h), and (2) thearrival times measured in steps j), l) and m), together with thetranslation distance of step k).